Confocal photoacoustic microscopy with optical lateral resolution

ABSTRACT

A confocal photoacoustic microscopy system includes a laser configured to emit a light pulse, a focusing assembly configured to receive the light pulse and to focus the light pulse into an area inside an object, an ultrasonic transducer configured to receive acoustic waves emitted by the object in response to the light pulse, and an electronic system configured to process the acoustic waves and to generate an image of the area inside the object. The focusing assembly is further configured to focus the light pulse on the object in such a way that a focal point of the focusing assembly coincides with a focal point of the at least one ultrasonic transducer.

STATEMENT REGARDING FEDERALLY SPONSORED RESEARCH & DEVELOPMENT

This invention was made with government support under grants R01EB000712 and R01 NS46214, both awarded by the U.S. National Institutesof Health. The government has certain rights in the invention.

CROSS REFERENCE TO RELATED APPLICATIONS

This application is a U.S. National Phase Patent Application ofInternational Application Serial Number PCT/US2008/081167 filed Oct. 24,2008, which is incorporated herein in its entirety, and which claimspriority from U.S. Provisional Patent Application 60/982,624 filed Oct.25, 2007, which is incorporated herein in its entirety.

BACKGROUND

The field of the invention relates generally to noninvasive imaging and,more particularly, to imaging an area with an object using confocalphotoacoustic imaging.

The capability of noninvasively imaging capillaries, the smallest bloodvessels, in vivo has long been desired by biologists at least because itprovides a window to study fundamental physiological phenomena, such asneurovascular coupling, on a microscopic level. Existing imagingmodalities, however, are unable to simultaneously provide sensitivity,contrast, and spatial resolution sufficient to noninvasively imagecapillaries.

BRIEF DESCRIPTION

In one aspect, a method for noninvasively imaging a scattering medium isprovided. The method includes focusing a light pulse into apredetermined area inside an object using a focusing assembly, andtransforming a photoacoustic signal into an acoustic wave, wherein thephotoacoustic signal is emitted by the object in response to the lightpulse. The method also includes detecting the acoustic wave using atransducer that is positioned such that the transducer and the focusingassembly are coaxial and confocal, and creating an image of thepredetermined area inside the object based on a signal generated by thetransducer representative of the acoustic wave.

In another aspect, a confocal photoacoustic imaging apparatus isprovided. The apparatus includes a focusing assembly configured toreceive a light pulse and to focus the light pulse into an area insidean object, a transducer configured to receive acoustic waves emitted bythe object in response to the light pulse, wherein the transducer ispositioned such that the transducer and the focusing assembly arecoaxial. The apparatus also includes a processor configured to recordand process the received acoustic waves.

In another aspect, a confocal photoacoustic microscopy system isprovided, including a laser configured to emit a light pulse, a focusingassembly configured to receive the at least one light pulse and to focusthe light pulse into an area inside an object, an ultrasonic transducerconfigured to receive acoustic waves emitted by the object in responseto the light pulse, and an electronic system configured to process theacoustic waves and to generate an image of the area inside the object.The focusing assembly is further configured to focus the light pulse onthe object in such a way that a focal point of the focusing assemblycoincides with a focal point of the ultrasonic transducer.

BRIEF DESCRIPTION OF THE DRAWINGS

Aspects of the invention may be better understood by referring to thefollowing description in conjunction with the accompanying drawings.

FIG. 1 is a diagram of a photoacoustic sensor that may be used with animaging system.

FIG. 2 is a block diagram of a system that uses confocal photoacousticmicroscopy.

FIG. 3 is a diagram of a photoacoustic sensor that may be used with theimaging system shown in FIG. 2.

FIG. 4 is a diagram of an alternative photoacoustic sensor that may beused with the imaging system shown in FIG. 2.

FIG. 5 is a diagram of a second alternative photoacoustic sensor thatmay be used with the imaging system shown in FIG. 2.

FIG. 6 is a schematic diagram of a third alternative photoacousticsensor that may be used with the imaging system shown in FIG. 2.

FIG. 7 is a schematic diagram of a fourth alternative photoacousticsensor that may be used with the imaging system shown in FIG. 2.

FIG. 8 is a diagram of a fifth alternative photoacoustic sensor that maybe used with the imaging system shown in FIG. 2.

FIG. 9 is a schematic diagram of a sixth alternative photoacousticsensor that may be used with the imaging system shown in FIG. 2.

FIG. 10 is a schematic diagram of a seventh alternative photoacousticsensor that may be used with the imaging system shown in FIG. 2.

FIG. 11 is a schematic diagram of an eighth photoacoustic sensor thatmay be used with the imaging system shown in FIG. 2.

FIGS. 12A-12C are images representing a lateral resolution measurementby the imaging system using a resolution test target immersed in clearliquid.

FIGS. 13A and 13B are images representing a measurement of the imagingdepth by the imaging system.

FIGS. 14A and 14B are photoacoustic images of a microvasculature by theimaging system.

FIG. 14C is a photograph of the microvasculature of FIGS. 14A and 14B,taken from a transmission microscope.

FIGS. 15A and 15B are maximum amplitude projection (MAP) images acquiredbefore and after a high-intensity laser treatment.

FIG. 16A is an in vivo image of a capillary bed captured using theimaging system.

FIG. 16B is an in vivo image of multiple levels of blood vesselbifurcations captured using the imaging system.

DETAILED DESCRIPTION

While the making and using of various embodiments of the invention arediscussed in detail below, it should be appreciated that the presentlydescribed embodiments provide many applicable inventive concepts thatmay be embodied in a wide variety of contexts. The embodiments discussedherein are merely illustrative of exemplary ways to make and useembodiments of the invention and do not delimit the scope of theinvention.

To facilitate the understanding of the presently described embodiments,a number of terms are defined below. Terms defined herein have meaningsas commonly understood by a person of ordinary skill in the areasrelevant to aspects of the invention. Terms such as “a,” “an,” “the,”and “said” are not intended to refer to only a singular entity, butinclude the general class of which a specific example may be used forillustration and are intended to mean that there are one or more of theelements. The terms “comprising,” “including,” and “having” are intendedto be inclusive and mean that there may be additional elements otherthan the listed elements.

The order of execution or performance of the operations in embodimentsof the invention illustrated and described herein is not essential,unless otherwise specified. That is, the operations may be performed inany order, unless otherwise specified, and embodiments of the inventionmay include additional or fewer operations than those disclosed herein.For example, it is contemplated that executing or performing aparticular operation before, contemporaneously with, or after anotheroperation is within the scope of aspects of the invention.

The terminology herein is used to describe embodiments of the invention,but their usage does not delimit the invention.

In embodiments of the invention, the terms used herein follow thedefinitions recommended by the Optical Society of America (OCIS codes).

In embodiments of the invention, the term “photoacoustic microscopy”includes, but is not limited to, a photoacoustic imaging technology thatdetects pressure waves generated by light absorption in the volume of amaterial, such as, but not limited to biological tissue, and propagatedto the surface of the material. Photoacoustic microscopy includes, butis not limited to, a method for obtaining images of the optical contrastof a material by detecting acoustic and/or pressure waves traveling froman object under investigation. Moreover, the term “photoacousticmicroscopy” includes, but is not limited to, detection of the pressurewaves that are still within the object.

In embodiments of the invention, the term “photoacoustic tomography”includes, but is not limited to, a photoacoustic imaging technology thatdetects acoustic and/or pressure waves generated by light absorption inthe volume of a material, such as, but not limited to biological tissue,and propagated to the surface of the material.

In embodiments of the invention, the term “piezoelectric detectors”includes, but is not limited to, detectors of acoustic waves utilizingthe principle of electric charge generation upon a change of volumewithin crystals subjected to a pressure wave.

In embodiments of the invention, the terms “reflection mode” and“transmission mode” includes, but is not limited to, a laserphotoacoustic microscopy system that employs the detection of acousticand/or pressure waves transmitted from a volume from which the waves aregenerated to an optically irradiated surface and a surface that isopposite to, or substantially different from, the irradiated surface,respectively.

In embodiments of the invention, the term “time-resolved detection”includes, but is not limited to, the recording of the time history of apressure wave with a temporal resolution sufficient to reconstruct thepressure wave profile.

In embodiments of the invention, the term “transducer array” includes,but is not limited to, an array of ultrasonic transducers.

In embodiments of the invention, the terms “focused ultrasonicdetector,” “focused ultrasonic transducer,” and “focused piezoelectrictransducer” include, but are not limited to, a curved ultrasonictransducer with a hemispherical surface or a planar ultrasonictransducer with an acoustic lens attached or an electronically focusedultrasonic array transducer.

In embodiments of the invention, the term “diffraction-limited focus”includes, but is not limited to, a best possible focusing of lightwithin limitations imposed by diffraction.

In embodiments of the invention, the term “confocal” includes, but isnot limited to, a situation when the focus of the illumination systemcoincides with the focus of the detection system.

The embodiments described herein relate to noninvasively imagingcapillaries. Some of the embodiments relate to microscopic photoacousticimaging using focused optical illumination and focused ultrasonicdetection. For example, an embodiment performs optical-resolutionphotoacoustic microscopy (OR-PAM), which facilitates providing a lateralresolution of 5 micrometers (μm) and a maximum imaging depth of greaterthan 0.7 millimeters (mm) based on endogenous optical absorptioncontrast. In vivo images of healthy capillary networks and lasercoagulated microvessels in mouse ears, for example, are demonstrated asexamples of applications of OR-PAM in biomedical research.

In an embodiment, the lateral resolution is dominantly determined by theoptical focus. A tightly focused optical illumination produces a localtemperature rise due to light absorption. The temperature rise leads tothermal expansion, which results in photoacoustic emission. Thephotoacoustic emission may be detected by a high-frequency largenumerical-aperture spherically focused ultrasonic transducer that iscoaxial and confocal with the light focusing system. The photoacousticemission may also be measured by an ultrasonic transducer array, a phasesensitive optical coherence tomography apparatus, a laser opticalinterferometer, and/or a capacitive surface displacement sensor. Byfocusing light to a focal spot of several micrometers in diameter,embodiments of the invention significantly improve the image resolutionof photoacoustic microscopy of biological tissue or other opticallyscattering media. It combines the high spatial resolution of opticalconfocal microscopy and the high optical absorption contrast ofphotoacoustic tomography.

The embodiments described herein provide for reflection-mode microscopicphotoacoustic imaging using focused optical illumination. Embodiments ofthe invention use a nearly diffraction-limited focused opticalillumination to achieve high spatial resolution. Embodiments of theinvention use a confocal arrangement between the optical focus and theultrasonic focus of a high-frequency large numerical-aperture (NA)spherically focused ultrasonic transducer to achieve high sensitivity.The ultrasonic transducer may be replaced with another detector capableof measuring local thermal expansion. By tightly focusing light, thelateral resolution limitations of existing photoacoustic microscopybased on the resolution of the ultrasonic focusing system may beovercome. In addition, because a photoacoustic signal is proportional tothe optical fluence at the target, the currently described embodimentsrequire only a low laser pulse energy and, hence, may be made relativelycompact, fast, and inexpensive. In the exemplary embodiment, a laserpulse energy of approximately 100 nanojoules (nJ) may be used.

Moreover, exemplary embodiments utilize optical focusing andtime-resolved detection of laser-induced pressure waves to obtainthree-dimensional images of the distribution of optical-absorptioncontrast within a sampling volume. The exemplary embodiments providenon-invasive imaging of scattering media, such as, but not limited to,biological tissue in vivo. The exemplary embodiments providenon-invasive imaging up to approximately one optical transport mean freepath deep. For most biological tissue, an optical transport mean freepath is approximately 1.0 millimeter (mm). In the exemplary embodiment,resolution on the order of 1.0 micrometer (μm) is attainable. Further,the exemplary embodiment images optical-absorption contrast inbiological tissue up to approximately 0.7 mm deep with a lateralresolution of approximately 5.0 μm. In embodiments of the invention, alarge numerical-aperture (NA) spherically focused ultrasonic transduceris used in a confocal coaxial arrangement with the light focusing opticsto facilitate providing high axial resolution of between 10.0 and 15.0μm.

An imaging procedure, which uses a confocal photoacoustic imagingsystem, is one of the possible embodiments and is aimed at medical andbiological applications. The presently described embodiments arecomplementary to the structural information that may be obtained frompure optical and ultrasonic imaging technologies and may be used fordiagnostic, monitoring or research purposes. Applications of thetechnology include, but are not limited to, the imaging of arteries,veins, capillaries (the smallest blood vessels), pigmented tumors suchas melanomas, and sebaceous glands in vivo in humans or animals. Thepresently described embodiments may use the spectral properties ofintrinsic optical contrast to monitor blood oxygenation (oxygensaturation of hemoglobin), blood volume (total hemoglobinconcentration), and even the metabolic rate of oxygen consumption; itmay also use the spectral properties of a variety of dyes or othercontrast agents to obtain additional functional or molecular-specificinformation. In other words, the presently described embodiments arecapable of functional and molecular imaging. Further, the presentlydescribed embodiments may be used to monitor possible tissue changesduring x-ray radiation therapy, chemotherapy, or other treatment. Inaddition, presently described embodiments may also be used to monitortopical application of cosmetics, skin creams, sun-blocks or other skintreatment products. The presently described embodiments, whenminiaturized, may also be used endoscopically, e.g., for the imaging ofatherosclerotic lesions in blood vessels.

Further, the presently described embodiments provide a method ofcharacterizing a target within a tissue by focusing one or more laserpulses on the region of interest in the tissue so as to penetrate thetissue and illuminate the region of interest, receiving the pressurewaves induced in the object by optical absorption using one or moreultrasonic transducers that are focused on the same region of interest,and recording the received acoustic waves so that the structure orcomposition of the object may be imaged. The one or more laser pulsesare focused by a microscope objective lens or a similar tightly focusingoptical system, which typically includes an optical assembly of lensesand/or mirrors, which converges the one or more laser pulses towards thefocal point of the ultrasonic transducer. The focusing device may alsouse one or more optical spatial filters, which may be a diaphragm or asingle-mode fiber, to reduce the focal spot of the optical system to thesmallest possible size so that the highest possible spatial resolutionmay be achieved. The focused one or more laser pulses selectively heatthe region of interest, causing the object to expand and produce apressure wave whose temporal profile reflects the optical absorption andthermo-mechanical properties of the target. Alternatively, an annulararray of ultrasonic transducers may be used along the tissue to enhancea depth of field of an imaging system by using synthetic aperture imagereconstruction. The signal recording includes digitizing the receivedacoustic waves and transferring the digitized acoustic waves to acomputer for analysis. The image of the object is formed from therecorded acoustic waves.

In addition, the presently described embodiments may also include anelectronic system in communication with the focusing device, the one ormore ultrasonic transducers, or a combination thereof. In oneembodiment, the electronic system includes an XYZ or circular scanner orscanners, an amplifier, a digitizer, a laser wavelength tuningelectronics, a computer, a processor, a display, a storage device, orcombination thereof. One or more component of the electronic system maybe in communication remotely with the other components of the electronicsystem, the apparatus, or both.

FIG. 1 shows a schematic of an exemplary focusing assembly 100 whichuses the confocal photoacoustic microscopy method. The light out of thedye laser is focused by a condenser lens 1 on a diaphragm (pinhole) 2for spatial filtering. Sampling beam splitter 10 is used to monitor thelaser output power through photo-detector 11 and to optically image theobject surface through eyepiece or aligning optics 12 for alignment. Thelight coming out of the spatial filter is focused by microscopeobjective lens 3 onto object 13 through beam separating element 6, 7, 9,and acoustic lens 8. Correction lens 5 placed on top of the beamseparation element compensates for the aberrations introduced by theprisms and the acoustic lens. The distance between the pinhole and theobjective lens is approximately 400 millimeters (mm), which gives anoptical focusing spot size of approximately 3.7 micrometers (μm) indiameter and a focal zone of approximately 200 μm in water. The laserpulse energy measured after the objective lens is approximately 100nanojoules (nJ). The beam separation element consists of an isoscelestriangular prism 6 with an apex angle of approximately 52.5° and arhomboidal 52.5° prism 7. Prisms 6 and 7 are adjoined along the diagonalsurfaces with a gap of approximately 0.1 mm in between. Gap 9 is filledwith an optical refractive-index-matching, low-acoustic-impedance,nonvolatile liquid such as 1000 cSt silicone oil, commercially availablefrom Clearco Products. The silicone oil and the glass have a goodoptical refractive index match (glass: 1.5; silicone oil: 1.4) but alarge acoustic impedance mismatch (glass: 12.1×106 N·s/m3; silicone oil:0.95×106 N·s/m3). As a result, the silicone oil layer is opticallytransparent but acted as an acoustic reflector. The photoacoustic signalemitted by the target is transformed by the acoustic lens, having aradius of curvature of approximately 5.2 mm, a diameter of approximately6.35 mm, a NA of approximately 0.46 in water, and an ultrasonic focalspot size of approximately 27 μm, into a plane elastic wave inrhomboidal prism 7 and is then detected by the high-frequencydirect-contact ultrasonic transducer 4 such as a model V2012-BCtransducer, commercially available from Panametrics-NDT with a centerfrequency of approximately 75 MHz, a bandwidth of approximately 80%, andan active element diameter of approximately 6.35 mm. Within thebandwidth of the ultrasonic transducer 4, ultrasonic absorption insilicone oil is high enough to dampen ultrasonic reverberations in thematching layer and thus minimize interferences to the image.

FIG. 2 is a block diagram of a system 200 based on confocalphotoacoustic microscopy, which is capable of contour scanning andquantitative spectroscopic measurement. The system includes a pulsedtunable laser 1 including a tunable laser pumped by a Q-switched laser,a focusing assembly 2, one or more ultrasonic transducers 4, and anelectronic system. The electronic system includes data acquisitionpersonal computer (PC) 3, motion controller 9, first and second scanners7 and 8, amplifier 5, and data acquisition subsystem (DAQ) 6, whichincludes a signal conditioner and a digitizer. Focusing assembly 2receives one or more laser pulses and focuses the one or more laserpulses into an area inside the sample object 10 so as to penetrate thetissue and illuminate the region of interest. The one or more ultrasonictransducers 4 are focused on the same the region of interest and receivethe acoustic or pressure waves induced in the region of interest by theone or more laser pulses. The electronic system records and processesthe received acoustic or pressure waves. The laser pulse generation,data acquisition, and object scanning are synchronized with the pulsesproduced by the motor controller at programmed locations of the laserfocus with respect to object 10. As described above, the focusingassembly 2 includes an optical assembly of lenses and/or mirrors thatfocuses one or more laser beams on the object in such a way that thefocal point of the optical focusing device coincides with that of theone or more ultrasonic transducers.

The focusing assembly is placed on an XYZ translation stage to performraster scanning along the object surface with simultaneous adjustment ofthe sensor's axial position to compensate for the curvature of theobject surface. Other embodiments may use different ways of imageformation, which include, but are not limited to, circular scanning,sector scanning, optical scanning, electronic focusing a transducerarray, and array-based image reconstruction. The recorded pressure-wavetime histories are displayed by the computer versus the focusingassembly position to construct a three dimensional image of thedistribution of the optical contrast within the tissue, i.e., a threedimensional tomographic image of the object.

System 200 employs a tunable dye laser, such as a model CBR-D laser,commercially available from Sirah, pumped by a neodymium-doped yttriumlithium fluoride (Nd:YLF) laser, such as the INNOSLAB laser,commercially available from Edgewave, as the irradiation source. Thelaser pulse duration is approximately 7 nanoseconds (ns) and the pulserepetition rate, which is controlled by the external triggering signal,is as high as approximately 2 kilohertz (kHz). In alternativeembodiments, a plurality of sources of penetrating radiation, which maybe confined to or concentrated in a small volume within the object, maybe used. Such sources include, but are not limited to, pulsed lasers,flash lamps, other pulsed electromagnetic sources, particle beams, ortheir intensity-modulated continuous-wave counterparts.

The one or more focused short laser pulses are delivered to an object(e.g., human or animal body, tissue or organ) under investigation, wherea small area of the object inside the focal area of the ultrasonictransducer is illuminated. The laser wavelength is selected as acompromise between the desired light penetration depth and the contrastbetween the structures of interest and the surrounding medium. Lightabsorption by the internal structures causes a transient temperaturerise which, due to thermoelastic expansion of the medium, produceselastic waves that may travel through the medium.

High-frequency ultrasonic waves generated in tissue by the laser pulseare recorded and analyzed by a PC to form a three-dimensional image. Theshape and dimensions of the optical-contrast structures are generallydetermined from the temporal profile of the laser-induced ultrasonicwaves and the position of the focusing assembly. Ordinarily, a rasterscan by the focusing assembly is used to form a three-dimensional image.However, a transducer array may be used to reduce the time of scanningand the total light exposure. When the tissue under investigation is aninternal organ, the optical fiber and ultrasonic transducer may beincorporated in an endoscope and positioned inside the body. Thefollowing examples will be provided for the purpose of illustratingvarious embodiments of the invention and are not meant to limit thepresent invention in any fashion.

As illustrated in FIGS. 3-6, the presently described embodiments providean optical resolution confocal microscopic photoacoustic imagingtechnology to image biological tissues in vivo. The exemplary embodimenthas a lateral resolution as high as approximately 5 μm and a maximumimaging depth of approximately 0.7 mm. In alternative embodiments, theimage resolution may be further improved by increasing the frequency ofthe ultrasonic transducer and the numerical aperture of the opticalobjective lens perhaps at the cost of imaging depth. The photoacousticimages shown in FIGS. 12-16 were obtained with minimal signal averagingand, therefore, could be further improved by averaging, at the expenseof data acquisition time, in another embodiment of the invention. Thecurrent imaging speed is limited by the pulse repetition rate of thelaser. Because lasers with pulse repetition rates of up to 100 KHz arenow available, other embodiments involve faster photoacoustic imaging,which can reduce motion artifacts, and extensive signal averaging.

The presently described embodiments include any realization of lightfocusing any kind of mirrors, lenses, fibers, and diaphragms that mayproduce well focused (preferably diffraction-limited) illuminationconfined to the focal area of the focused ultrasonic transducer. Thepresently described embodiments also cover any confocal photoacoustictechniques with any light delivery and detection arrangements in whichthe lateral resolution is defined by the focusing of the incidentradiation rather than the acoustic detection unit.

One or more of the following embodiments may be used to implement laserfocusing for the purpose described herein: (1) an optical microscopeobjective lens that focuses a well-collimated single-mode lased beaminto a nearly diffraction-limited point, (2) an objective lens thatforms an image of a small pinhole on the region of interest, (3) afocusing system in which a single-mode optical fiber is used instead ofpinhole, (4) a focusing system in which an oscillating mirror scans theoptical focus rapidly within the larger focal area of the ultrasonictransducer. The following embodiments, and further alternativeembodiments, may also be used to implement laser focusing for further,undescribed purposes. Various examples of the focusing assembly will nowbe described in reference to FIGS. 3-10, wherein the focusing assemblyincludes, for example, an optical focusing device, and one or moreultrasonic transducers in the piezoelectric, optical, or another form.

FIG. 3 is a diagram of a focusing assembly 300 of imaging system 200(shown in FIG. 2). A custom-made cubic beam splitter or right-angleprism 4 with a sub-micron reflective aluminum coating layer 6 sandwichedbetween the two prisms is used to couple the optical and ultrasonicradiations. A pair of optical objective lenses 1 focuses the laser lightfrom the single-mode optical fiber onto the region of interest insidethe object, where metal coating 6 is used to reflect the optical beam. Asampling beam splitter 8 is placed between the objective lenses 1 tomonitor the laser output power with a photo-detector 9 and to view theobject surface for alignment with an eyepiece or aligning optics 10.Ultrasonic radiation emitted by the object 11 passes through an acousticlens 5, the aluminum optical reflector, and reaches an ultrasonictransducer 2.

FIG. 4 is a diagram of a focusing assembly 400 of imaging system 200(shown in FIG. 2). A laser pulse from a pulse laser is focused by acondenser lens 1 on a diaphragm 2 for spatial filtering. The lightcoming out of the spatial filter 2 is reflected by an oscillating mirror10, which performs fast optical scanning within the wider focal area ofan ultrasonic transducer 4. The laser beam is focused into an object bya microscope objective lens 3 through a beam splitting element 6, 7, 9and an acoustic lens 8. A thin plano-convex optical lens 5 is placed ontop of the beam splitting element 6, 7, 9 to compensate for theaberrations introduced by the prisms 6 and 7 and the acoustic lens 8.The beam splitting element 6, 7, 9 consists of an isosceles triangularprism 6 with an apex angle of 52.5° and a rhomboidal 52.5° prism 7.Prisms 6 and 7 are adjoined along the diagonal surfaces but areseparated by a thin layer of refractive-index-matching,low-acoustic-impedance, and nonvolatile liquid, such as alow-molecular-weight silicone oil 9. The photoacoustic signal emitted bythe object is transformed by the acoustic lens 8 into a plane elasticwave in rhomboidal prism 7. Ultrasonic reflection from the boundary ofsilicone oil 9 converts at least 98% of the energy of the incidentlongitudinal wave into that of a shear wave, which is transformed backinto a longitudinal wave on the free surface of rhomboidal prism 7 andthen detected by high-frequency direct-contact ultrasonic transducer 4.Because the acoustic focus is generally several times wider than theoptical focus, taking advantage of fast optical scanning in thisembodiment may significantly decrease the image acquisition time.

FIG. 5 is a diagram of a focusing assembly 500 of imaging system 200(shown in FIG. 2). An optical objective lens 2 focuses the outputaperture of a single-mode optical fiber 1 into the object through theoptically clear slit window in a one-dimensional ultrasonic arraytransducer 4 placed on an optically transparent substrate 5. Substrate 5serves as a wave-guide for acoustic waves and may have a cylindricalfocus acoustic lens on its outer surface. The light coming out of thespatial filter is reflected by an oscillating mirror 3, which performsfast optical scanning. Ultrasonic radiation emitted by the object iscollected by ultrasonic transducer array 4. A multiple-elementpiezoelectric transducer array may accelerate the image acquisition timein one dimension owing to the electronic focusing of the transducerarray. The acoustic focus provided by assembly 500 follows the focalposition of the laser beam without mechanically scanning the ultrasonictransducer over the object. Three-dimensional images may be acquired bymechanically translating the focusing assembly perpendicularly to theslit.

FIG. 6 is a diagram of a focusing assembly 600 of imaging system 200(shown in FIG. 2). The light output from a single-mode optical fiber 1is reflected by a mirror scanner 2, collimated by an optical objectiveor excitation lens 3, passed through a dichroic mirror 4, and thenfocused by another objective lens 5 on a region of interest through aFabry-Perot etalon 6, which is acoustically coupled to the object.Mirror scanner 2 performs rapid 2D raster scanning of the object bysweeping the excitation laser beam. The photoacoustic wave from theobject causes a transient strain distribution in Fabry-Perot etalon 6,which shifts its resonance wavelengths. Another laser (probing laser) 9working at a different optical wavelength scans over Fabry-Perot etalon6 through a second mirror scanner 8, a second objective lens 7, anddichroic mirror 4 to read the strain distribution in Fabry-Perot etalon6. The strain is then converted into the photoacoustic pressuredistribution. In the exemplary embodiment, no mechanical scanning isnecessary to form a 3D image of the object.

FIG. 7 is a diagram of a focusing assembly 700 of imaging system 200(shown in FIG. 2). An optical objective lens 4 focuses the outputaperture of a single-mode optical fiber 1 into a region of interest inan object to excite photoacoustic waves. A 2D mirror scanner 3 isintroduced in the optical path to perform 2D scanning of the object. Aphase-sensitive optical coherence tomography (OCT) system 5 working at adifferent optical wavelength is focused on the same region of interestby the optical objective lens 4 and 2D mirror scanner 3. The two lightbeams of different wavelengths are coupled by a dichroic mirror 2. Thephase-sensitive OCT system measures, within the optical focal spotinside the object, the photothermal effect due to absorption of thelaser pulse. The photothermal effect in the object is measured beforepressure waves propagate to the surface of the object. In the exemplaryembodiment, focusing assembly 700 forms a 3D image without translatingthe objective lens and does not require direct contact with the object.Correspondingly, it may be potentially very fast and may be used wherenon-contact imaging is preferred.

FIG. 8 is a diagram of an alternative embodiment of the focusingassembly 800 suitable for hand-held operation. An optical objective lens4 images the aperture of a single-mode optical fiber 1 onto the regionof interest in the object through an optically clear window in aspherically focused ultrasonic transducer 5. A sampling beam splitter 2reflects a small portion of the incident light to monitor the laseroutput power with a photo-detector 3. The ultrasonic radiation emittedby the object is received by the ultrasonic transducer 5. Thephotoacoustic assembly is mounted on a pendulum 6, which is attached toa frame 8 through a flexible mount, such as a flat spring 7. The frameis water-tight and contains optically transparent acoustic couplingfluid, such as water, for light delivery and acoustic coupling. Moved byan actuator 9, pendulum 6 may perform sector scanning of the objectrapidly. A position sensor 10 monitors the position of the optical focusand is used to synchronize the pulse laser so that image distortion dueto varying scanning velocity is minimized.

FIG. 9 is a diagram of another alternative embodiment of a focusingassembly 900 suitable for applications inside body cavities such asinter-vascular imaging. A laser pulse delivered by a single-mode fiber 1is focused on the region of interest in the object by an optical lensassembly 4 through an optically clear window in a spherically focusedultrasonic transducer 6. Ultrasonic transducer 6 together with aright-angled prism 5 is connected to a flexible shaft 2 located inside acatheter 3. Optically and acoustically transparent circular window 7allows the optical beam and ultrasonic radiation to pass freely to andfrom the object. Photoacoustic images are formed by rotating the shaft 2with respect to the axis of the catheter and axially translating thecatheter.

FIG. 10 is a block diagram of another alternative embodiment of afocusing assembly 1000 which uses the confocal photoacoustic microscopymethod simultaneously with optical confocal microscopy. The light comingout of the pulsed laser is focused by a condenser lens 1 on a diaphragm(pinhole) 2 for spatial filtering. A dichroic beam splitter or mirror 10is used to monitor the laser output power with a photo-detector 11 andto form an optical fluorescence confocal image of the object. Theoptical fluorescence confocal imaging portion consists of a pinhole, ordiaphragm, 12, a focusing system or lens 13, a low pass optical filter15, and a photo-detector (such as a photomultiplier) 14. The lightcoming out of the spatial filter is focused by a microscope objectivelens 3 on the object through a beam splitting element. The beamsplitting element consists of an isosceles triangular prism 6 with anapex angle of 52.5° and a rhomboidal 52.5° prism 7. Prisms 6 and 7 areadjoined along the diagonal surfaces with a gap in between 9. Gap 9 isfilled with refractive-index-matching, low-acoustic-impedance,nonvolatile liquid. A correction lens 5 is placed on top of the beamsplitting element to compensate for aberrations introduced by the prismsand the acoustic lens. The photoacoustic signal emitted by the object istransformed by an acoustic lens 8 into a plane elastic wave inrhomboidal prism 7. Ultrasonic reflection from the boundary of the prismconverts the incident longitudinal elastic wave into a shear wave. Theshear wave propagates toward the free surface of the rhomboidal prism,where it is transformed back into a longitudinal wave and detected by ahigh-frequency direct-contact ultrasonic transducer 4 for imageformation and spectral measurements of the target.

The fusion of the optical confocal microscopy and photoacousticmicroscopy provides complementary information about the object. Onefeature is the quantitative measurement of the optical absorptionspectrum of the object by simultaneously using the fluorescence signalfrom the optical confocal microscope and the photoacoustic signal fromthe photoacoustic microscope. The quantitative measurement of theoptical absorption spectrum of the object requires knowledge of thespectral variation of the excitation optical fluence at the focus, whichmay be measured using the fluorescent signals as illustrated below.

In the exemplary embodiment, two excitation optical wavelengths areused. If a fluorescence dye is present, the detected fluorescence signalV_(f)(λ_(xi),λ_(m)) at the i-th excitation wavelength λ_(xi) and theemission wavelength λ_(m) is a product of the unknown local excitationoptical fluence, the concentration of dye C, the known molar opticalabsorption coefficient of the dye ε_(af) (λ_(xi)), the quantum yield ofthe dye Q, and the fluorescence detection sensitivity S_(f)(λ_(m)). Fori=1 and 2, the following ratio in Equation (1) is present:

$\begin{matrix}{\frac{V_{f}\left( {\lambda_{x\; 1},\lambda_{m}} \right)}{V_{f}\left( {\lambda_{x\; 2},\lambda_{m}} \right)} = {\frac{{ɛ_{af}\left( \lambda_{x\; 1} \right)}{F\left( \lambda_{x\; 1} \right)}}{{ɛ_{af}\left( \lambda_{x\; 2} \right)}{F\left( \lambda_{x\; 2} \right)}}.}} & (1)\end{matrix}$

Therefore, the local excitation optical fluence ratio may be recoveredas in Equation (2):

$\begin{matrix}{\frac{F\left( \lambda_{x\; 1} \right)}{F\left( \lambda_{x\; 2} \right)} = {\frac{V_{f}\left( {\lambda_{x\; 1},\lambda_{m}} \right)}{V_{f}\left( {\lambda_{x\; 2},\lambda_{m}} \right)}/{\frac{ɛ_{af}\left( \lambda_{x\; 1} \right)}{ɛ_{af}\left( \lambda_{x\; 2} \right)}.}}} & (2)\end{matrix}$

Similarly, the detected photoacoustic signal V_(pa)(λ_(xi)) is a productof the local excitation optical fluence F(λ_(xi)), the opticalabsorption coefficient of dominantly absorbing hemoglobin μ_(ah)(λ_(xi))and the acoustic detection sensitivity Sa. Assuming that the hemoglobinabsorbs much more than the fluorescent dye, the following ratio inEquation (3) is developed:

$\begin{matrix}{\frac{V_{pa}\left( {\lambda_{x\; 1},\lambda_{m}} \right)}{V_{pa}\left( {\lambda_{x\; 2},\lambda_{m}} \right)} = {\frac{{\mu_{ah}\left( \lambda_{x\; 1} \right)}{F\left( \lambda_{x\; 1} \right)}}{{\mu_{ah}\left( \lambda_{x\; 2} \right)}{F\left( \lambda_{x\; 2} \right)}}.}} & (3)\end{matrix}$

From the above two equations, the ratio of the hemoglobin absorptioncoefficient may be recovered as in Equation (4):

$\begin{matrix}{\frac{\mu_{ah}\left( \lambda_{x\; 1} \right)}{\mu_{ah}\left( \lambda_{x\; 2} \right)} = {\frac{V_{pa}\left( {\lambda_{x\; 1},\lambda_{m}} \right)}{V_{pa}\left( {\lambda_{x\; 2},\lambda_{m}} \right)}\frac{V_{f}\left( {\lambda_{x\; 2},\lambda_{m}} \right)}{V_{f}\left( {\lambda_{x\; 1},\lambda_{m}} \right)}{\frac{ɛ_{af}\left( \lambda_{x\; 1} \right)}{ɛ_{af}\left( \lambda_{x\; 2} \right)}.}}} & (4)\end{matrix}$

This ratio may be used to quantify the oxygen saturation of hemoglobinand the relative total concentration of hemoglobin. Of course, thisexample merely illustrates the principle, which may be extended to themeasurement of other optical absorbers using two or more excitationoptical wavelengths.

The presently described embodiments may be used to estimate oxygenmetabolism in tissues and organs, by combining measurements of bloodflow and oxygenation into and out of regions of interest. Oxygenmetabolic rate (MRO2) is the amount of oxygen consumed in a given tissueregion per unit time per 100 grams (g) of tissue or of the organ ofinterest. Since in typical physiological conditions, hemoglobin is thedominant carrier of oxygen, the key measure of blood oxygenation isoxygen saturation of hemoglobin (SO2), as follows in Equation (5):MRO₂∝(SO_(2,in)—SO_(2,out))·C_(Hb)·A_(in)· v _(in).  (5)

Here, A_(in) is the area of the incoming vessel, v _(in) is the meanflow velocity of blood in the incoming vessel, and C_(Hb) is the totalconcentration of hemoglobin. While A_(in) and v _(in) may be estimatedusing ultrasound imaging, SO2 and relative C_(Hb) may be estimated frommulti-wavelength photoacoustic methods.

Exemplary advantages of photoacoustic microscopy over traditionaloptical and ultrasonic imaging include the detection of endogenousoptical absorption contrast at ultrasonic resolution. In photoacousticmicroscopy, a pulsed laser beam is focused into the biological tissue toproduce emission of ultrasonic waves due to the photoacoustic effect.The short-wavelength pulsed ultrasonic waves are then detected with afocused ultrasonic transducer to form high-resolution tomographicimages. Among the existing photoacoustic imaging technologies, thespatial resolutions depend almost solely on the ultrasonic parametersincluding the center frequency, bandwidth, and numerical aperture (NA).For example, using dark-field confocal PAM, a lateral resolution ofapproximately 50 μm has been achieved with a center frequency ofapproximately 50 megahertz (MHz) and an NA of approximately 0.44. Thisresolution from prior systems is inadequate to resolve smallerstructures such as capillaries between approximately 3 μm andapproximately 7 μm in diameter with endogenous optical absorptioncontrast. Aspects of the invention provide improved spatial resolution.

If such an improvement is achieved by increasing the ultrasonic focusingcapability, an approximately 5-μm lateral resolution requires anultrasonic center frequency greater than 300 MHz. At such a highfrequency, unfortunately, the ultrasonic attenuation, which isapproximately 400 μm⁻¹ in water and 100 μm⁻¹ in tissue, limits thepenetration depth to approximately 100 μm. An alternative is to use fineoptical focusing to provide the lateral resolution while ultrasonictemporal detection provides axial resolution. Such an alternative,called OR-PAM, is primarily sensitive to optical absorption contrast,whereas conventional reflection-mode optical confocal microscopy isdominantly sensitive to scattering or fluorescence.

FIG. 11 is a schematic of another exemplary embodiment of the OR-PAMimaging system. In this embodiment, the system employs nearlydiffraction limited optical focusing with bright field opticalillumination to achieve μm-level lateral resolution. A dye laser, suchas a CBR-D laser commercially available from Sirah, pumped by a Nd:YLFlaser is used as the irradiation source. The laser pulse duration isapproximately 5 ns and the pulse repetition rate, controlled by anexternal trigger, is as high as 2 kHz. The light from the dye laser isattenuated by one thousand times, passed through a spatial filter, suchas a 25 μm pinhole, commercially available as P250S from Thorlabs, andthen focused by a microscope objective lens, such as a RMS4X lensavailable from Thorlabs and including a NA of approximately 0.1, a focallength of approximately 45 mm, and a working distance of approximately22 mm. The distance between the pinhole and the objective lens isapproximately 400 mm. The input aperture of the microscope objective isapproximately 0.8 times the diameter of the Airy disk of the spatialfilter. As a result, the diffraction-limited focus of the objective inwater is approximately 3.7 μm in diameter and approximately 200 μm infocal zone. The laser pulse energy after the objective lens measuresapproximately 100 nJ. An optional beam splitter is located between thepinhole and the objective lens to facilitate focus adjustment and systemalignment. Two right-angled prisms, the NT32-545 prism available fromEdmund Optics, for example, form a cube with a gap of approximately 0.1mm between the hypotenuses. The gap is filled with silicone oil. Asdescribed above, the silicone oil and the glass have a good opticalrefractive index match but a large acoustic impedance mismatch. As aresult, this silicone oil layer is optically transparent butacoustically reflecting. An ultrasonic transducer, such as a V2012-BCtransducer available from Panametrics-NDT, with a center frequency of 75MHz, a bandwidth of 80%, and an active-element diameter of 6.35 mm, isattached to a cathetus of the bottom prism as shown in FIG. 11. Aplano-concave lens with an approximately 5.2 mm radius of curvature andan approximately 6.35 mm aperture is attached to the bottom of the cubeto function as an acoustic lens, which has an NA of approximately 0.46in water and a focal diameter of approximately 27 μm. Of course, thislens also functions as a negative optical lens, which is compensated forby a correcting positive optical lens placed on top of the cube.

The photoacoustic signal detected by the ultrasonic transducer isamplified by approximately 48 dB using, for example, two ZFL 500LNamplifiers commercially available from Mini-Circuits, then digitized bya 14-bit digital acquisition board using, for example, a CompuScope12400 from Gage Applied Sciences. A raster scanning is controlled by aseparate PC, which triggers both the data-acquisition PC and the pumplaser. The trigger signal is synchronized with the clock-out signal fromthe digital acquisition board.

An acoustic lens is immersed in water inside a heated container. Awindow is opened at the bottom of the container and sealed with anultrasonically and optically transparent 25-μm thick polyethylenemembrane. The animal is placed under the water tank with the region ofinterest (ROI) exposed below the window. Ultrasonic gel, such as ClearImage, available from SonoTech, is applied to the ROI for acousticcoupling. For simplicity, the raster scanning is implemented bytranslating the water tank and the animal together along the horizontal(x-y) plane. One-dimensional (1D) photoacoustic signal (A-line) at eachhorizontal location is recorded for 1 μs at a sampling rate of 200 MS/s.A volumetric photoacoustic image is formed by combining thetime-resolved photoacoustic signals and may be viewed in directvolumetric rendering, cross-sectional (B-scan) images, or maximumamplitude projection (MAP) images.

FIGS. 12A-12C are images representing a lateral resolution measurementby the imaging system. FIG. 12A is a MAP image of an Air Forceresolution test target, FIG. 12B is a magnified image of the regionwithin the dashed box of FIG. 12A, and FIG. 12C is a MAP image of a6-μm-diameter carbon fiber. The lateral resolution of the OR-PAM systemwas experimentally measured by imaging an Air Force resolution testtarget immersed in clear liquid. Images were acquired at the opticalwavelength of approximately 590 nm and no signal averaging was performedduring data acquisition. In FIGS. 12A and 12B, the highlightedwell-resolved bars, shown as group 6, element 5, have gaps ofapproximately 4.9 μm, a spatial frequency of approximately 102 mm⁻¹, anda modulation transfer function value of 0.65. Other pairs of spatialfrequency and modulation transfer function values include, for example,a 64 mm⁻¹ spatial frequency with a 0.95 modulation transfer functionvalue, and a 80 mm⁻¹ spatial frequency with a 0.8 modulation transferfunction value. Nonlinearly fitting of the modulation transfer functionyields a lateral resolution of approximately 5 μm, which is 30% greaterthan the diffraction limit of 3.7 μm. As an illustration of the lateralresolution, an MAP image of a 6-μm-diameter carbon fiber immersed inwater is shown in FIG. 12C. The mean full-width-at-half-maximum (FWHM)value of the imaged fiber is approximately 9.8 μm, which is 3.8 μm widerthan the fiber diameter and hence in agreement with the ˜5 μmresolution. The axial resolution was estimated to be approximately 15 μmbased on the measured transducer bandwidth, approximately 100 MHz inreceiving-only mode, and the speed of sound in tissue, approximately 1.5mm/μs. In tissue, both the lateral and the axial resolutions deterioratewith imaging depth because of optical scattering and frequency-dependentacoustic attenuation, respectively.

FIGS. 13A and 13B are images representing a measurement of the imagingdepth by the imaging system. FIG. 13A is a MAP image of two horse hairsplaced above and below a piece of rat skin acquired with the OR-PAMsystem. FIG. 13B is a B-scan image at the location marked by the dashedline in FIG. 13A. The imaging depth of this system was measured byimaging two horse hairs with a diameter of approximately 200 μm placedabove and below a piece of freshly harvested rat scalp. A photoacousticimage was acquired with 32 times signal averaging at the opticalwavelength of 630 nm. Both hairs are clearly visible, where the bottomhair shows a weaker photoacoustic signal because of both optical andacoustic attenuation in the skin. The B-scan image shows that the bottomhair is 700 μm deep in the tissue. Therefore, the maximum imaging depthis at least 700 μm.

The microvessels in the ear of a nude mouse were imaged in vivo by thisOR-PAM at the optical wavelength of 570 nm. Nude mouse ears having athickness of approximately 300 μm have well-developed vasculature andhave been widely used to study tumor angiogenesis and othermicrovascular diseases. During image acquisition, the animal was keptmotionless using a breathing anesthesia system and was kept warm usingan infrared lamp. Unlike studies published elsewhere, no opticalclearing agent was applied to the skin surface. An area of 1 mm² wasscanned with a step size of approximately 1.25 μm. For each pixel, 16(i.e., 4 by 4) neighboring A-lines were averaged to increase thesignal-to-noise ratio (SNR). The scanning time for a complete volumetricdataset was approximately 18 minutes. After data acquisition, the animalrecovered naturally without observable laser damage.

FIGS. 14A and 14B are photoacoustic images of a microvasculature by theimaging system. FIG. 14C is a photograph of the microvasculature ofFIGS. 14A and 14B, taken from a transmission microscope. Morespecifically, FIG. 14A is an in vivo photoacoustic image ofmicrovasculature in a nude mouse ear, FIG. 14B is a 3D visualization ofthe volumetric photoacoustic data with pseudocolor, and FIG. 14C is aphotograph taken with trans-illumination optical microscopy. In FIGS.14A-14C, the area denoted as C is a capillary, the area denoted as CB isa capillary bed, and the area denoted as SG is a sebaceous gland. Thephotoacoustic image of the microvasculature (FIGS. 14A and 14B) agreeswith the photograph (FIG. 14C) taken from a transmission microscope witha 4× magnification. However, capillaries are imaged by only the OR-PAMsystem described above. The mean ratio of the photoacoustic amplitudesbetween the blood vessels and the background is 20:1, which demonstratesa high endogenous optical-absorption-based imaging contrast. Somevessels, e.g. the vessel labeled with C in FIG. 14A, only occupy asingle pixel, which presumably indicates a capillary with a diameter ofapproximately 5 μm. A volumetric rendering of the photoacoustic data(FIG. 14B) shows the three-dimensional connectivity of the bloodvessels. Parallel arteriole-venule pairs and their branching are clearlyobserved. The diameter and the morphological pattern of the vesselwithin the dashed-box in both FIGS. 14A and 14B suggest that thesemicrovessels belong to a capillary bed. Therefore, OR-PAM, as describedabove, is able to image capillaries in vivo with endogenous opticalabsorption contrast due to hemoglobin. In addition, sebaceous glands mayalso be imaged at the same time.

FIGS. 15A and 15B are MAP images acquired before and after ahigh-intensity laser treatment. FIG. 15A is an in vivo photoacousticimage of laser-induced vessel destruction in a Swiss Webster mouse earbefore a laser treatment. FIG. 15B is an in vivo photoacoustic imageafter the laser treatment. In FIGS. 15A and 15B, the area denoted as Tis the laser treated area, the area denoted as W is widened bloodvessels, and the area denoted as H is a possible hemorrhage. To furtherdemonstrate the potential of OR-PAM, the high-intensity laserdestruction of microvessels in the ear of a Swiss Webster mouse wereimaged. This type of destruction is clinically used to remove port winestains in humans. FIGS. 15A and 15B show the MAP images acquired beforeand after the high-intensity laser treatment. After the healthyvasculature was imaged by the OR-PAM system (shown in FIG. 15A), thecenter region measuring approximately 0.25×0.25 mm² was treated byhigh-intensity laser pulses having a peak optical fluence ofapproximately 10 J/cm² scanned with the step size of approximately 1.25μm. For the high-intensity illumination, the attenuator and the pin holewere removed from the light path. A second image (shown in FIG. 15B) wasacquired 15 minutes after the laser treatment. Disruption of the vesselswithin the treated region was clearly observed in the dashed box.Further, the destruction of the blood vessels dilated severalneighboring vessels and produced possibly hemorrhage.

FIG. 16A is an in vivo photoacoustic image of a capillary bed in a mouseear, captured using the OR-PAM imaging system with a focusing depth ofapproximately 50 μm. FIG. 16B is an in vivo photoacoustic image ofmultiple levels of blood vessel bifurcations in a mouse ear, capturedusing the OR-PAM imaging systems with a focusing depths of approximately150 μm.

The embodiments described herein use (1) optical focusing to achievehigh lateral resolution, (2) time-resolved detection of laser-inducedpressure waves to obtain high axial resolution, and/or (3) confocalarrangement between the optical excitation and ultrasonic receiving focito achieve high sensitivity. In alternative embodiments, the focusedultrasonic receiving may be replaced with optical sensing of thephotothermal effect directly inside the object. Three-dimensional imagesof the distribution of optical contrast within a sampling volume areacquired.

In an existing system, an intensity-modulated continuous-wave beam ofradiation is combined with the detection of the magnitude of thephotoacoustic signal. In the embodiments described herein, short pulsedexcitation is combined with time-resolved detection of the photoacousticsignal, which has the advantage of time-of-flight based axialresolution. Therefore, the presently described embodiments provide, forexample, (a) enhanced axial resolution, (b) 3D imaging of opticalcontrast from a 2D raster scan, and (c) minimal image artifacts due tothe interference of photoacoustic waves from various targets within thelight illumination volume, in contrast to the existing system.

Another existing system uses focused light to produce thermal expansionand uses optical detection, based on the thermal lens effect, or anultrasonic detector to monitor the resulting pressure/densitytransients. Such a system lacks axial resolution. In addition, thelateral resolution of such a system is determined by the detector ratherthen the excitation optics. Utilization of the thermal lens effect insuch a system requires transmission illumination in an optically clearmedium, which limits the applications of the technology. Moreover, inusing an unfocused ultrasonic transducer and an unfocused ultrasonicdetector, the excitation beam has a large separation, which affects thedetection sensitivity. The frequency mismatch between the centralfrequency of the photoacoustic waves (>100 MHz) and the centralfrequency of the ultrasonic transducer (<10 MHz) also limits the SNR ofsuch a system.

Another existing system uses laser excitation in a coaxial arrangementwith a focused ultrasonic detection. However, the laser beam used insuch a system is not focused. In fact, the laser beam is divergentbecause the positive acoustic lens functions as a negative optical lens.The negative optical lens actually broadens the optical beam. Moreimportantly, such a system neither achieves nor claims optically definedlateral resolution, which is a key feature in the presently describedembodiments.

The ability to image microstructures such as the micro-vascular networkin the skin or brain cortex and monitor physiological functions oftissue is invaluable. One of the promising technologies foraccomplishing this objective is photoacoustic microscopy. Currenthigh-resolution optical imaging techniques, such as optical coherencetomography, can image up to approximately one transport mean free pathof between 1 to 2 mm into biological tissues. However, such techniquesare sensitive to the backscattering that is related to tissuemorphology, and are insensitive to the optical absorption that isrelated to important biochemical information. Other known techniquessuch as confocal microscopy and multi-photon microscopy have even morerestrictive penetration depth limitation and often involve theintroduction of exogenous dyes, which with a few notable exceptions haverelatively high toxicity. Acoustic microscopic imaging and spectroscopysystems are sensitive to acoustic impedance variations, which havelittle functional information about biological tissue and have lowcontrast in soft tissue. Other imaging techniques such as diffuseoptical tomography or thermal wave microscopy have low depth toresolution ratio. Photoacoustic imaging as in embodiments of theinvention provides high optical-absorption contrast while maintaininghigh penetration depth and high ultrasonic resolution. Moreover, becausephotoacoustic wave magnitude is, within certain bounds, linearlyproportional to the optical contrast, optical spectral measurement canbe performed to gain functional, i.e., physiological, information suchas the local blood oxygenation level. However, increasing the resolutionpower beyond several tens of micrometers meets serious challenges. Atultrasonic frequencies required to achieve such resolution, which isabove approximately 100 MHz, ultrasonic absorption in tissue graduallybecomes proportional to the square of the ultrasonic frequency.Consequently, a resolution of several micrometers will have penetrationdepth of a few tens of micrometers that is much less than thepenetration depth of other optical imaging techniques such as confocalmicroscopy. Embodiments of the present invention overcome the resolutionlimitation by using optical focusing to achieve high lateral resolutionand ultrasonic detection to achieve axial resolution.

Although imaging of photothermal treatment of microvessels itself isbiomedically significant, the capability of OR-PAM to imagephysiological and pathological changes in capillaries has broaderapplications. Other possible applications include imaging ofvasodilation and vasoconstriction in stroke models, tumor angiogenesis,and tumor extravasations. Mouse ears were chosen as the initial organ totest OR-PAM because transmission optical microscopy could be used tovalidate the photoacoustic images. Since OR-PAM operates in reflectionmode, it may be applied to many other anatomical sites.

Several alternative embodiments are possible. First, photoacousticimages may be acquired by scanning the optical-acoustic dual fociinstead of the sample and the transducer container. Second, it ispossible to scan only the optical focus within the acoustic focusingarea to reduce the image acquisition time significantly. Third, byvarying the excitation optical wavelength, physiological parameters suchas hemoglobin oxygen saturation and blood volume may be quantified forin vivo functional imaging using endogenous contrast. Similarly,targeted exogenous contrast agents such as indocyanine green (ICG) andnanoparticles may be quantified for in vivo molecular imaging. Fourth,the acoustic coupling cube may be made to transmit photoacoustic wavesten times more efficiently without transformation from p-waves intosv-waves so that the SNR may be improved. Acoustic antireflectioncoating on the lens should further increase the SNR by approximately 10dB.

When the optical focus is 100 μm below the tissue surface, the surfaceoptical fluence is close to the ANSI safety limit of 20 mJ/cm2 in thevisible spectral region. Although the ANSI standards regulate only thesurface fluence, the spatial peak optical fluence is calculated at thefocus in water, which is approximately 500 mJ/cm2. This focal fluence isstill less than the experimentally observed damage threshold in livetissue. After the aforementioned improvements are implemented, theoptical fluence may be reduced without affecting the SNR.

It will be understood that particular embodiments described herein areshown by way of illustration and not as limitations of the invention.The principal features in embodiments of this invention may be employedin various embodiments without departing from the scope of theinvention. Those skilled in the art will recognize, or be able toascertain using no more than routine experimentation, numerousequivalents to the procedures described herein. Such equivalents areconsidered to be within the scope of this invention and are covered bythe claims.

All of the compositions and/or methods disclosed and claimed herein maybe made and executed without undue experimentation in light of thepresent disclosure. While embodiments of this invention have beendescribed in terms of preferred embodiments, it will be apparent tothose of skill in the art that variations may be applied to thecompositions and/or methods and in the steps or in the sequence of stepsof the method described herein without departing from the concept,spirit, and scope of the invention. All such similar substitutes andmodifications apparent to those skilled in the art are deemed to bewithin the spirit, scope and concept of the invention as defined by theappended claims.

It will be understood by those of skill in the art that information andsignals may be represented using any of a variety of differenttechnologies and techniques (e.g., data, instructions, commands,information, signals, bits, symbols, and chips may be represented byvoltages, currents, electromagnetic waves, magnetic fields or particles,optical fields or particles, or any combination thereof). Likewise, thevarious illustrative logical blocks, modules, circuits, and algorithmsteps described herein may be implemented as electronic hardware,computer software, or combinations of both, depending on the applicationand functionality. Moreover, the various logical blocks, modules, andcircuits described herein may be implemented or performed with a generalpurpose processor (e.g., microprocessor, conventional processor,controller, microcontroller, state machine, and/or combination ofcomputing devices), a digital signal processor (“DSP”), an applicationspecific integrated circuit (“ASIC”), a field programmable gate array(“FPGA”), or other programmable logic device, discrete gate ortransistor logic, discrete hardware components, or any combinationthereof designed to perform the functions described herein. Similarly,steps of a method or process described herein may be embodied directlyin hardware, in a software module executed by a processor, or in acombination of the two. A software module may reside in RAM memory,flash memory, ROM memory, EPROM memory, EEPROM memory, registers, harddisk, a removable disk, a CD-ROM, or any other form of storage mediumknown in the art. Although embodiments of the present invention havebeen described in detail, it will be understood by those skilled in theart that various modifications may be made therein without departingfrom the spirit and scope of the invention as set forth in the appendedclaims.

This written description uses examples to disclose the invention,including the best mode, and also to enable any person skilled in theart to practice the invention, including making and using any devices orsystems and performing any incorporated methods. The patentable scope ofthe invention is defined by the claims, and may include other examplesthat occur to those skilled in the art. Such other examples are intendedto be within the scope of the claims if they have structural elementsthat do not differ from the literal language of the claims, or if theyinclude equivalent structural elements with insubstantial differencesfrom the literal languages of the claims.

What is claimed is:
 1. A method for noninvasively imaging a scatteringmedium, said method comprising: focusing at least one light pulse into apredetermined area inside an object using a focusing assembly;transforming absorbed optical energy into an acoustic wave, the acousticwave emitted by the object in response to the light pulse; channelingthe acoustic wave through a beam separation element that is eitheracoustically reflective and optically transparent or opticallyreflective and acoustically transparent; detecting the acoustic waveusing at least one transducer positioned such that the transducer andthe focusing assembly are coaxial and confocal; and creating an image ofthe predetermined area inside the object based on a signal generated bythe transducer representative of the acoustic wave.
 2. A methodaccording to claim 1, wherein focusing at least one light pulsecomprises focusing the at least one light pulse using a focusing opticalsystem such that the at least one light pulse converges at a focal pointof the transducer.
 3. A method according to claim 2, wherein focusing atleast one light pulse further comprises reducing the focal point of thefocusing optical system to facilitate increasing a spatial resolution.4. A method according to claim 1, wherein creating an image comprisesrecording and digitizing the received acoustic wave so that thepredetermined area inside the object may be imaged.
 5. A confocalphotoacoustic imaging apparatus comprising: a focusing assemblyconfigured to receive at least one light pulse and to focus the lightpulse into an area inside an object, said focusing assembly comprising abeam separation element that is either acoustically reflective andoptically transparent or optically reflective and acousticallytransparent; at least one transducer configured to receive acousticwaves emitted by the object in response to the light pulse, said atleast one transducer positioned such that said at least one transducerand said focusing assembly are coaxial; and a processor configured torecord and process the received acoustic waves.
 6. A confocalphotoacoustic imaging apparatus according to claim 5, wherein saidfocusing assembly comprises an optical assembly of lenses and mirrorsconfigured to focus the at least one light pulse on the object in such away that a focal point of said focusing assembly coincides with a focalpoint of said at least one ultrasonic transducer.
 7. A confocalphotoacoustic imaging apparatus according to claim 5, wherein saidfocusing assembly focusing assembly is positioned on an XYZ translationstage in order to perform raster scanning along a surface of the objectwith simultaneous adjustment of an axial position of said apparatus tocompensate for a curvature of the surface of the object.
 8. A confocalphotoacoustic imaging apparatus according to claim 5, wherein saidfocusing assembly comprises an optical microscope objective lensconfigured to focus the at least one light pulse into a nearlydiffraction-limited point.
 9. A confocal photoacoustic imaging apparatusaccording to claim 5, wherein said focusing assembly comprises anobjective lens configured to form an image of a small pinhole on thearea inside the object.
 10. A confocal photoacoustic imaging apparatusaccording to claim 5, wherein said focusing assembly comprises anoscillating mirror configured to scan an optical focus rapidly within alarger focal area than a focal area of said at least one transducer. 11.A confocal photoacoustic microscopy system comprising: a laserconfigured to emit at least one light pulse; a focusing assemblyconfigured to receive the at least one light pulse and to focus thelight pulse into an area inside an object, said focusing assemblycomprising a beam separation element that is either acousticallyreflective and optically transparent or optically reflective andacoustically transparent; at least one ultrasonic transducer configuredto receive acoustic waves emitted by the object in response to the lightpulse, wherein said focusing assembly is further configured to focus theat least one light pulse on the object in such a way that a focal pointof said focusing assembly coincides with a focal point of said at leastone ultrasonic transducer; and an electronic system configured toprocess the acoustic waves and to generate an image of the area insidethe object.
 12. A confocal photoacoustic microscopy system according toclaim 11, wherein said focusing assembly focusing assembly is positionedon an XYZ translation stage in order to perform raster scanning along asurface of the object with simultaneous adjustment of an axial positionof said apparatus to compensate for a curvature of the surface of theobject.
 13. A confocal photoacoustic microscopy system according toclaim 11, wherein said focusing assembly comprises an optical microscopeobjective lens configured to focus the at least one light pulse into anearly diffraction-limited point.
 14. A confocal photoacousticmicroscopy system according to claim 11, wherein said focusing assemblycomprises an objective lens configured to form an image of a smallpinhole on the area inside the object.
 15. A confocal photoacousticmicroscopy system according to claim 11, wherein said focusing assemblycomprises a single-mode optical fiber configured to emit a focused atleast one light pulse.
 16. A confocal photoacoustic microscopy systemaccording to claim 11, wherein said focusing assembly comprises anoscillating mirror configured to scan an optical focus rapidly within alarger focal area than a focal area of said at least one transducer. 17.A confocal photoacoustic microscopy system according to claim 11,wherein said electronic system is configured to record the acousticwaves and to display the recorded acoustic waves versus a correspondingposition of said focusing assembly in order to generate the image.
 18. Aconfocal photoacoustic microscopy system according to claim 17, whereinsaid electronic system comprises a motor controller configured tosynchronize data acquisition and object scanning with the at least onelight pulse at programmed locations of the laser focus with respect tothe object.